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Experimental System Prototype of a Portable, Low-Cost, C-Scan Ultrasound Imaging Device Michael I. Fuller*, Member, IEEE, Karthik Ranganathan, Member, IEEE, Shiwei Zhou, Member, IEEE, Travis N. Blalock, Member, IEEE, John A. Hossack, Senior Member, IEEE, and William F. Walker, Member, IEEE
Abstract—A system prototype of a future compact, low-cost medical ultrasound device is described and presented with experimental results. The prototype system consists of a 32 32 element, fully sampled 2-D transducer array and a printed circuit board (PCB) containing 16 custom “front-end” receive channel integrated circuits (ICs) with analog multiplexing and programmable logic. A PC that included a commercially available data acquisition card is used for data collection and analysis. Beamforming is performed offline using the direct sampled in-phase/quadrature (DSIQ) algorithm. Pulse-echo images obtained with the prototype are presented. Results from this prototype support the feasibility of a low-cost, pocket-sized, C-scan imaging device. Index Terms—Application-specific integrated circuits, biomedical acoustic imaging, C-scan, portable ultrasound.
I. INTRODUCTION
A
S DIGITAL beamforming techniques are refined and advances in semiconductor technology continue to allow for smaller, more power-efficient, and less expensive integrated circuits (ICs), the ultrasound community has begun focusing on developing portable medical ultrasound devices [1]–[11]. These compact systems allow for the expansion of ultrasound into application areas formerly excluded by system size and cost considerations. The newer systems are lightweight (often “hand-held”) units that better facilitate patient point-of-care and offer an order of magnitude cost reduction. Examples of these newer, smaller commercial systems include: the iLook series (Sonosite, Inc., Bothell, WA); OptiGo (Philips Medical Systems, Andover, MA); Acuson Cypress (Siemens Medical Solutions USA, Malvern, PA); HS-1500 (Honda Electronics, Co., Ltd., Toyohashi, Aichi, Japan); Terason t3000 (Teratech
Manuscript received February 17, 2007; revised May 16, 2007. This work was supported by the Carilion Biomedical Institute and NIH NIBIB Grant RO1 EB0023489. Asterisk indicates corresponding author. *M. I. Fuller was with the Department of Biomedical Engineering, University of Virginia, Charlottesville, VA 22908 USA. He is now with PocketSonics, Inc., Charlottesville, VA 22901 USA (e-mail:
[email protected]). K. Ranganathan was with the Department of Biomedical Engineering, University of Virginia, Charlottesville, VA 22908 USA. He is now with PocketSonics, Inc., Charlottesville, VA 22901 USA. S. Zhou was with the Department of Biomedical Engineering, University of Virginia, Charlottesville, VA USA. He is now with Philips Research North America, Briarcliff Manor, NY 10510 USA. T. N. Blalock is with the Department of Electrical and Computer Engineering, University of Virginia, Charlottesville, VA 22904 USA. J. A. Hossack is with the Department of Biomedical Engineering, University of Virginia, Charlottesville, VA 22908 USA. W. F. Walker is with the Department of Biomedical Engineering and the Department of Electrical and Computer Engineering, University of Virginia, Charlottesville, VA 22904 USA. Digital Object Identifier 10.1109/TBME.2007.903517
Corporation, Burlington, MA); the LOGIQ Book (GE Medical Systems, Waukesha, WI); and the z.one (ZONARE Medical Systems, Inc., Mountain View, CA). While these devices have already found widespread use, a vast niche for medical ultrasound remains unfilled among unconventional or uninitiated users of ultrasound. Such clinicians, medical technicians, battlefield medics, and veterinarians would greatly benefit from using ultrasound to provide adjunct information during routine medical examinations or when rapid diagnosis is crucial to patient survival. Adapting the development of a medical ultrasound device to these clinical needs requires judicious tradeoffs in system complexity to provide an avenue to creating a pocket-sized unit with a simple, intuitive interface and order of magnitude cost reduction beyond that offered by current hand-held systems. Furthermore, by employing a 2-D transducer array, more intuitive scan formats such as C-mode become available to users who may have had limited exposure to B-mode imaging. The Sonic Window is a low-cost, pocket-sized medical ultrasound system currently under development at the University of Virginia with these concepts and target applications in mind. Potential applications for the Sonic Window include guiding needle and catheter insertion [12]–[16]; guiding biopsies [17], [18]; locating foreign bodies [19], [20]; identifying internal bleeding and fluid collection [13], [21]; and supporting routine physical examination [3], [21]–[24]. The low cost and compact size of the device could open applications in veterinary medicine and animal research. One such application would be tumor localization and growth monitoring, a task that is currently performed via palpation or visual inspection. In this and other applications, the Sonic Window is envisioned as a widely available, easy to use device suitable for users with little or no prior ultrasound experience. As such, it will not compete with state of the art systems with respect to image quality or flexibility, but must surpass them with respect to ease of use, cost, and portability. The high level of integration required by the Sonic Window requires a significant degree of collaboration and cross-innovation between transducer, electronics, and beamforming development. There exists a need, therefore, to experimentally verify the expected performance and tradeoffs of these three components not only in isolation, but also as a system, and in a manner versatile enough to accommodate modifications in their design. It is probable that several versions of the transducer will be designed and fabricated as the dicing procedure, backing design, and material selection are refined. Also, any modifications to the receive electronics circuitry will require new ICs to be fabricated.
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Fig. 2. System block diagram of the proposed Sonic Window device. The design is partitioned into a single transmit circuit, a fully sampled 2-D transducer array, a custom IC containing receive circuitry, a commercially available DSP, and a liquid crystal display.
Fig. 1. Long-term concept of the Sonic Window device. A model of the Sonic Window is used to demonstrate its use in guiding needle insertion, one of many possible applications.
Equally important, the development of an experimental system containing a custom, fully sampled 2-D array and custom receive circuitry provides a unique opportunity to study the performance of a variety of beamforming and signal processing algorithms while having knowledge and control over design parameters that are inaccessible on commercial clinical scanners. This paper discusses the design and assembly of an experimental ultrasound system consisting of a 32 32-element, fully sampled 2-D transducer array, a PCB containing sixteen 64-channel front-end receive ICs with associated control logic and bias circuits, a PCI card digitizer, and a PC that executes beamforming and signal processing algorithms in software. The 2-D transducer array was fabricated on a separate PCB and mates with the receive electronics PCB by way of surface mount connectors. This experimental system can capture and store 1024 channels of RF data for subsequent offline processing, enabling us to explore the effects of our transducer fabrication, circuit topologies, and beamforming strategies on image quality. A description of the system and obtained experimental results are provided next.
A block diagram of the future Sonic Window system is illustrated in Fig. 2. The design is partitioned into a single transmit circuit, a fully-sampled 2-D transducer array, a custom IC containing an array of receive channels, a commercially available DSP, and a liquid crystal display. The common node of the 2-D array is connected to the transmit circuit during transmit mode and to analog ground during receive mode. Each receive channel will consist of an on-chip transmit protection shunting device, a variable gain preamplifier, a bandpass filter, a sampleand-hold (S/H) stage, an analog-to-digital converter (ADC), and static memory. Placing the transmit protection devices on-chip eliminates the need for bulky, expensive, and power-consuming off-chip switching elements. The S/H stage consists of two S/H units, whose output samples are one quarter period apart at the center frequency of the received pulse. This operation estimates the in-phase (I) and quadrature (Q) components of the RF signal, as defined by the DSIQ beamforming algorithm [27]. Furthermore, by forming C-mode images rather than B-mode, each S/H unit need only capture a minimum of one sample per image. The combination of these two properties dramatically simplifies the design of the ADC by permitting digitization rates as low as 10 kHz and produces much less stringent memory and data bandwidth requirements. Since a standard CMOS process is used, this results in significant reductions in cost, IC area, and power consumption. III. DESCRIPTION OF EXPERIMENTAL SYSTEM PROTOTYPE
II. SYSTEM-LEVEL DESCRIPTION A description of the proposed ultracompact, low-cost medical ultrasound device, the Sonic Window, was presented earlier [25], [26]. The ultimate concept (Fig. 1) is a fully integrated, pocket-sized unit consisting of a 2-D array, receive and protection circuitry implemented on a custom integrated circuit, beamforming implemented on a digital signal processor (DSP), and a high-resolution LCD screen for image display. Crucial savings in circuit area and complexity are gained through the use of a novel beamforming approach, direct sampled in-phase/quadrature (DSIQ) beamforming [27], as well as new integrated circuit topologies [25], [35]. These innovations, combined with the development of an inexpensive method for fabricating a fully sampled 2-D transducer array on a PCB [28], enable an unprecedented level of integration and dramatic reductions in system cost.
An experimental ultrasound system was designed and constructed (Fig. 3) consisting of a 2-D array, custom receive and protection circuitry, and a PC that included a Gage Compuscope 12100 PCI card (Gage Applied Technologies, Inc., Lachine, QC, Canada). The transducer array was fabricated on a PCB substrate [28]. Each transducer element was electrically connected to a dedicated pad on a surface mount connector on the back of the PCB, while all the elements shared a common in in connection on the top. A separate ten-layer, PCB was fabricated (Fig. 4) that contains 16 custom receive circuitry ICs, four ADG707 8:1 differential analog multiplexing ICs (Analog Devices, Norwood, MA), two output buffer channels, and an onboard XCR3064XL complex programmable logic device (CPLD) (Xilinx, Inc., San Jose, CA). The PCB containing the 2-D transducer array was designed to mate with the receive circuitry PCB by way of the surface-mount
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The PCB (MetroCircuits, Rochester, NY) consists of nine routing layers to fan out each element pad to a contact on one of 16 surface-mount (SM) connectors on the rear of the board [Fig. 5(b)], which electrically mate the transducer array to corresponding SM connectors on the receive circuitry PCB. A tenth layer serves as a ground plane. Design parameters such as trace width and spacing (3 mil/76 m), minimum drill size (6 mil/152 m), and pad size (12 mil/305 m) represent the threshold of current industry capabilities for a ten-layer PCB. B. Custom-Integrated Circuit Design
Fig. 3. Schematic diagram of the experimental prototype system. The transducer array was fabricated on a PCB substrate. A separate PCB contains 16 custom receive circuitry ICs each consisting of 64 front-end receive channels. Each transducer element is electrically connected to a unique receive channel on the electronics PCB. The outputs of the 1024 receive channels are multiplexed on-chip (64:2) and off-chip (32:2) down to two bandpass filter output channels coupled to a PC fitted with a two-channel data acquisition card.
connectors (FX11-series, Hirose Electric (USA), Inc., Simi Valley, CA). The data acquisition card was capable of capturing two channels of 12-bit data simultaneously at a sampling rate of 50 MHz. Thus, 512 separate transmit events were necessary to capture all 1024 channels of RF data, and memory depth in each channel was limited to 1600 12-bit samples. Volume RF data acquired with the prototype system was sampled and stored on a PC for subsequent processing and data analysis in Matlab (The MathWorks, Inc., Natick, MA). A. Fully Sampled, 2-D Transducer Array Transducer development for the Sonic Window [28] involves the dual challenge of constructing a fully sampled (high channel count) 2-D array that is also inexpensive. The largest difficulty is the interconnect, where an electrical connection must be made between each transducer element and its dedicated receive channel. To date, we have used a printed circuit board substrate, which can be fabricated much less expensively than other approaches such as multilayer flex circuits, wire-bonding, and solder connections. The transducer was fabricated by attaching a 0.53-mm-thick wafer of high permittivity, high-electromechanical-coupling-coefficient ceramic (HD3203, CTS Wireless, Albuquerque, NM) to an array of pads on the PCB using Chobond silver epoxy (Chomerics, Woburn, MA). The wafer was then diced using a NBC-ZH2040 Disco dicing blade (Disco Corp., Tokyo, Japan) to form isolated elements. The kerfs were 0.04 mm wide and 0.125 mm deep into the PCB substrate. A low-viscosity, unfilled epoxy (RE2039, Loctite, Rocky Hill, CT) was used to fill the kerfs. Finally, the entire transducer array was covered with a gold-plated polyester electrode that serves as a common node to all the elements [Fig. 5(a)].
A number of researchers have recently implemented frontend receive electronics on custom ICs for applications such as real-time 3-D imaging [29], intravascular ultrasound [30], [31], intra-oral ultrasound [32], high-frequency annular arrays [33], and portable ultrasound [6], [8], [34]. We designed a custom IC containing 64 analog front-end receive channels implemented in a standard TSMC 0.35- m CMOS process available through the MOSIS Integrated Circuit Fabrication Service (Marina del Rey, CA). Sixteen of these 64-channel ICs are used to form all 1024 receive channels in our prototype system. Each channel consists of an on-chip transmit protection shunting device, a variable gain preamplifier, and a transconductance buffer. The receive channel is fully differential to reduce distortion and suppresses the effects of power supply and substrate noise. The transmit protection scheme, described in [35], excludes the expensive and area-consuming off-chip components used in other systems to prevent the high-voltage transmit pulse from damaging the receive electronics. Instead, a suitably sized NMOS transistor implemented on-chip is connected between the preamplifier input and a low-impedance power supply serving as analog ground. During the transmit event, this NMOS transistor is turned on, shunting the large current transient from the high-voltage transmit pulse to analog ground. Only a fraction (on the order of 100 mV) of the transmit voltage (as high as 100 V) appears at the input to the preamplifier. The shunt device is turned off during receive to permit amplification of the received echo signal. The low-noise preamplifier design (Fig. 6) consists of two identical differential stages with variable gain. The gain can be adjusted between 30 and 85 dB by adjusting the bias voltage of triode-region device M3, which serves as a source degeneration resistance (Fig. 7). A novel low-frequency suppression scheme is incorporated into the preamplifier design to serve the dual purpose of reducing 1/f noise and rejecting dc offset. The active load of the amplifier is designed to have high mid-band gain, but small low-frequency gain, as shown in Fig. 7. The gain profile can be tuned by adjusting the bias voltage LowFreqAdj. The preamplifier equivalent input noise of 5 nV/ Hz (within the band of interest) was found in simulation by performing a noise analysis in the Cadence Virtuoso Spectre Circuit Simulator (Cadence Design Systems, Inc., San Jose, CA). Each preamplifier is followed by a differential transconductance buffer performing a voltage-to-current conversion, providing the means to implement a current-mode analog multiplexing scheme in which multiple channels can share the same output node with minimal impact on signal bandwidth. The 64 channels are grouped into two 32-channel banks each having a
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Fig. 4. Photograph of the 11 in x 11.5 in receive electronics PCB (top view). The PCB contains sixteen 64-channel custom receive circuitry ICs, four differential analog multiplexing ICs, two output buffer channels, and an onboard CPLD. The 2-D transducer array was designed to mate with the rear side of the receive circuitry PCB by way of surface-mount connectors.
differential output (Fig. 8). One channel per bank is permitted to drive its signal onto the common output nodes at any given time. The output buffers of the inactive channels are disabled by turning off their respective output MOSFET devices, effectively forcing them into a high-impedance state. A 5-bit decoder selects the active channel in each bank. The active die area of the front-end receive circuit IC is 1.9 mm 0.9 mm, and includes the 64 analog receive channels and associated multiplexing logic (Fig. 9). The die were packaged (Promex Industries, Santa Clara, CA) into a Kyocera PGA121M (Kyocera America, Inc., San Diego, CA) 121-pin ceramic package. IV. METHODS The experimental prototype system was assembled and tested for basic functionality prior to attempting a transmit event or forming pulse-echo images. Test points on the back of the transducer array were probed with the transducer PCB connected to the receive electronics PCB. A 100-mV sinusoidal test signal was applied to each element and RF data were acquired from
all channels. This experiment provided verification of the mapping of transducer elements to channels and identified open or shorted connections between the transducer elements and the inputs to the receive circuitry ICs. Pulse-echo volume datasets were acquired using the prototype system. The transducer was driven at its center frequency of 3.3 MHz with an eight-cycle Gaussian-enveloped sinusoid having a full-width at half-maximum (FWHM) bandwidth of approximately 30% and a peak-to-peak amplitude of 30 V. As described above, the data acquisition hardware was capable of acquiring and storing 1024 channels of data sampled at 50 MHz with a memory depth of 1600 12-bit samples. C-mode images of each target were formed on the PC in Matlab using three beamforming techniques: beamforming using only time delays, conventional baseband demodulated I/Q beamforming using only phase delays, and DSIQ beamforming that also uses only phase delays. Beamforming was implemented to be consistent with the methods followed in [27], with the exception of the assumed transducer fractional bandwidth (30% versus 55%), center frequency (3.3 MHz versus 5.5 MHz), the sampling rate (50 MHz versus 39.27 MHz), and that fact that our experimental system
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Fig. 6. Schematic of a single stage of the differential preamplifier. The gain can be adjusted from 30 to 85 dB by adjusting the bias voltage RecvGain. The active load of the amplifier is designed to have high mid-band gain, but small low-frequency gain. The gain profile can be tuned by adjusting the bias voltage LowFreqAdj.
Fig. 5. Fully sampled 2-D transducer array PCB. The transducer array is located on the top of the PCB (a) where the gold foil common node is visible. The surface-mount connectors on the bottom of the PCB (b) interface with the receive electronics PCB and connect each transducer element to its respective receive channel.
forms C-mode images (in keeping with the Sonic Window device concept) as opposed to the B-mode images formed in [27]. -dB fracThe acquired data was filtered in Matlab to 30% tional bandwidth at 3.3 MHz with a 51-order FIR filter. Hann window apodization was used [36]. Scaling was applied where applicable to compensate for energy differences in pixels beamformed with receive apertures that intersected edges of the 2-D array. Time-delay beamforming was implemented using unquantized time delays (IEEE double precision floating point representation). A cubic spline-based continuous representation of the sampled data was used to evaluate the received signals in each channel at the requisite time points [27], [37]. RF data produced by summing across channels were envelope detected using the Hilbert transform.
Fig. 7. Experimental measurements of gain (at 5 MHz) as a function of RecvGain control voltage and frequency response of preamplifier at maximum gain setting (RecvGain = 2 V).
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Fig. 8. Block diagram of the 64-channel receive circuitry front-end IC. Each channel is fully differential and consists of an NFET protection device, a preamplifier, and a transconductance output buffer. The channels are arranged in two 32-channel banks. All channels within a bank share a differential output node. A 5-bit decoder selects one channel per bank at a time by enabling the channel’s output buffer.
As described in [27], DSIQ samples will be formed in our final system by splitting the received signal between two parallel sample-and-hold (S/H) channels. As described in Section II, the Sonic Window device will ultimately include S/H stages and ADCs on the same IC as the front-end receive circuitry components. In each receive channel, one S/H will acquire the I component and then the other S/H will acquire the Q component by sampling one quarter of a period later at the center frequency of the signal. In the experimental system prototype, only the front-end circuitry was implemented on-chip, and the sampling was performed by the Gage Compuscope 12100 PCI card at a rate of 50 MHz at uniform sampling intervals. The I component sample was taken directly from this acquired RF data—the Q component sample was synthesized by interpolating the acquired RF data at a time lag of a quarter period at the assumed center frequency using cubic spline interpolation [27], [37]. The apodization, phase rotation and summation across channels were implemented as in the conventional baseband demodulation case described above. The three beamforming methods described above were used to form C-mode images from pulse-echo data acquired off two targets for the purpose of comparing the performance of the DSIQ method to the time-delay and conventional I/Q methods, as well as evaluate the overall performance of the experimental system. The first target was a 200- m nylon wire in a water tank placed 1.5 cm below and parallel to the face of the transducer. The second target was a custom-made “edge phantom,” m/s) having which consisted of 10% acrylamide gel ( one speckle-generating region and one nonspeckle-generating region. The speckle-generating region was constructed by incorporating Sephadex (Amersham, Piscataway, NJ) into the acrylamide. The procedure followed in constructing this phantom was based on that described in [38], although higher acrylamide concentrations were used. The geometry of this edge phantom is such that an ideal C-mode image acquisition should produce an image with speckle in one half and an absence of speckle (anechoic region) in the other half. V. EXPERIMENTAL RESULTS
Fig. 9. Microphotograph of the 64-channel receive circuitry IC [35].
Conventional baseband demodulated data were obtained to form by multiplying the received data by to form the the in-phase (I) component and by quadrature (Q) component, where was the transducer center frequency of 3.3 MHz, was the sample number, and was the sampling interval. This data were then low-pass filtered with a fifth-order Butterworth filter and zero-phase distortion was accomplished by filtering once in the forward direction and then filtering a second time after reversing the output. This -dB cutoff at 3.3 MHz. The I and Q components produced a were combined into an analytic representation of the received echo signal that was then apodized and focused via phase rotation through complex multiplication operations, the results of which were summed across channels to yield the intensity at a given point on the image plane.
Fig. 10 is a binary mapping of “dead” channels discovered from the probing procedure. It was found that 69 out of 1024 electrical channels (6.74%) contained an open circuit between the transducer element and its corresponding receive channel input. The cause of the majority of these open connections was determined to be due to errors in the layout design of the 2-D transducer array PCB. Poor contact was also noted between the surface-mount connectors on the transducer PCB and the receive electronics PCB. Since contact between these surface-mount connectors relies on a friction fit, minor bending and warping of the transducer PCB can result in localized misalignment of connector contacts. This overall net channel yield should be distinguished from the transducer element yield, which was measured to be 99%. The top panel in Fig. 11 illustrates the simulated 2-D point spread function (PSF) in the C-mode plane for the ideal case in which 100% of the receive channels are connected to their respective transducer elements. The bottom panel in Fig. 11 illustrates the simulated 2-D PSF in the C-mode plane for a 6.74%
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Fig. 10. Diagram mapping the location of dead channels (defined as an open circuit between the receive channel and transducer element, indicated in black) with respect to their corresponding physical location on the 2-D transducer array.
channel loss having the same distribution (channel-to-element mapping) observed in the experimental prototype system (see Fig. 10). These two PSFs were computed in Matlab using a narrowband Rayleigh–Sommerfeld formulation [39] for a plane parallel to the transducer at a depth of 2 cm. Hann window apodization [37] was used. The images in Fig. 11 were normalized, logarithmically compressed, and mapped so as to present to 0 dB. image intensity over the range The C-mode images of the wire target acquired with the prototype system (Fig. 12) were formed using conventional unquantized time-delay beamforming (TD), conventional baseband demodulated I/Q beamforming (IQ), and DSIQ beamforming (DSIQ) methods for f#s of 0.5, 1, and 2. Hann window apodization was used. The C-mode images were normalized, logarithmically compressed, and mapped so as to 0 dB. The to present image intensity over the range FWHM of the wire in the acquired images at f/1 is 1.2 mm. One-dimensional integrated cross-sections of the wire target images in azimuth were formed by performing a 2-D spline interpolation (8x) on the original absolute image, summing in elevation, and then plotting the normalized, logarithmically compressed result. C-mode images of the edge phantom target acquired with the prototype system (Fig. 13) were formed using pure time-delay beamforming (TD), conventional baseband demodulated I/Q beamforming (IQ), and DSIQ beamforming (DSIQ) methods for f#s of 0.5, 1, and 2. Hann window apodization was used. The contrast between the region with speckle and region with no speckle was 10 dB. The C-mode images were normalized, logarithmically compressed and mapped so as to present image to 0 dB. One-dimensional inteintensity over the range grated cross-sections of the edge phantom images in elevation were formed by performing a 2-D spline interpolation (8x) on
Fig. 11. Simulated 2-D PSF (C-mode plane at a depth of 2 cm) for the ideal case in which 100% of the receive channels are connected to their respective transducer elements, and for a 6.74% channel loss following the distribution (channel-to-element mapping) observed in the experimental prototype system. Both images were normalized, logarithmically compressed and mapped so as to present image intensity over the range 40 to 0 dB.
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the original absolute image, summing in azimuth, and then plotting the normalized, logarithmically compressed result. VI. DISCUSSION Overall, the experimental results shown are very promising in that they demonstrate successful pulse-echo image formation with a low-cost, compact, experimental ultrasound system in its proof-of-concept phase of development. There do exist, however, multiple avenues for improving image quality. The largest contributor to poor image quality was the presence of large grating lobes (observable in Fig. 11 approximately 1.5 cm from the focus) caused by the transducer element pitch. The pitch was 635 m, while the wavelength at 3.3 MHz and 1540 m/s speed of sound is 467 m. Element pitch was limited by the PCB manufacturing capabilities, specifically the trace width/spacing and minimum pad size for vias. The minimum pad size was necessarily increased to account for the worsening drill tolerance as the aspect ratio (ratio of PCB thickness to
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Fig. 12. C-mode images acquired with the prototype system of a 200-m nylon wire in a water tank placed 1.5 cm below and parallel to the face of the transducer. The images were formed using pure time-delay beamforming (TD), conventional baseband demodulated I/Q beamforming (IQ), and DSIQ beamforming (DSIQ) methods for f#s of 0.5, 1, and 2. Hann window apodization was used. The images were normalized, logarithmically compressed, and mapped so as to present image intensity over the range 30 to 0 dB. One-dimensional integrated cross-sections of the wire target images in azimuth were formed by performing a 2-D spline interpolation (8x) on the original absolute image, summing in elevation, and then plotting the normalized, logarithmically compressed result.
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drill hole size) increased. The aspect ratio increased with the number of routing layers, which in turn depended on the trace width/spacing. Another significant contributor to image quality limitations is the manner in which signals were routed from individual transducer elements to their respective electronic receive channels. As described in Section V, routing errors in the transducer PCB along with intermittent poor contact between the surface mount connectors resulted in a channel loss of 6.74%. As illustrated in Fig. 11, these lost channels have the undesirable effect of significantly degrading the PSF by raising side lobe levels to as high as dB and distorting the mainlobe. Note also that the channel loss results in a shift variant system—pixel-to-pixel gain varies as the receive aperture is translated across this nonuniform pattern of “dead” transducer elements. These effects significantly contribute to the presence of clutter and other artifacts, causing anomalous spots and kinks or gaps in the wire target image and poor contrast in the edge phantom image. Furthermore, the inter-
connect scheme also suffers from parasitic inductances, capacitances, and resistances associated with long PCB traces, the surface-mount connectors, and the custom IC pin grid array (PGA) chip packages that all contribute to signal-to-noise (SNR) degradation, crosstalk, and impedance variation among channels. Despite successful image formation using this preamplifier (see Fig. 6), a significant problem was discovered involving the low-frequency suppression scheme. When a large transient voltage appeared at the input of the preamplifier (similar to that accompanying a transmit event), the drain-source voltage of the triode-region device M8 (M9) was large enough that a charge was drained from the MOScap, M12 (M13), in the active load. This changed the bias point enough to significantly lower the overall mid-band gain of the preamplifier. A temporary solution was implemented in which the LowFreqAdj bias was dynamically tuned such that after each transmit event some of the charge was allowed to return onto M12 (M13), though the gain and dynamic range of the overall preamplifier was still degraded.
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Fig. 13. C-mode images acquired with the prototype system of a custom-made “edge phantom,” which consisted of 10% acrylamide gel (c = 1545 m/s) having one speckle-generating region and one nonspeckle-generating region. The geometry of this edge phantom is such that an ideal C-mode image acquisition should produce an image with speckle in the bottom half and an absence of speckle (anechoic region) in the top half. The images were formed using pure time-delay beamforming (TD), conventional baseband demodulated I/Q beamforming (IQ), and DSIQ beamforming (DSIQ) methods for f#s of 0.5, 1, and 2. Hann window apodization was used. The images were normalized, logarithmically compressed, and mapped so as to present image intensity over the range 20 to 0 dB. One-dimensional integrated cross sections of the edge phantom images in elevation were formed by performing a 2-D spline interpolation (8x) on the original absolute image, summing in azimuth, and then plotting the normalized, logarithmically compressed result.
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Another factor influencing image quality is the use of the DSIQ beamforming algorithm, which makes concessions in image quality to provide the dramatic hardware savings exploited by the Sonic Window device. However, as illustrated in Figs. 12 and 13, the C-mode images formed using DSIQ beamforming exhibit very little difference from the conventional time-delay and I/Q methods. While it can be argued that the transducer, interconnect, and front-end shortcomings described above dominate the subtler differences in image quality that exist between the three beamforming approaches, the integrated 1-D cross sections in Figs. 12 and 13 offer further insight. Here, the differences between beamforming approaches are more evident—particularly at lower f#s—but the overall performance is still similar. DSIQ and conventional I/Q beamforming appear to behave almost identically, and both deviate from TD beamforming at lower f#s because both
approaches rely on phase rotation to achieve delays and are thus susceptible to focusing errors toward the edge of the aperture. The integrated 1-D cross-sections at higher f#s demonstrate closer agreement among all the beamforming approaches. These trends are similar to the findings in [27], where the DSIQ algorithm was shown to compare favorably with conventional beamforming techniques in a B-mode commercial ultrasound scanner not suffering from the grating lobe and channel-loss issues present in our experimental system. DSIQ beamforming was also shown to be surprisingly robust to error in the assumed center frequency, which would cause the S/H clocks to be offset at phase differences other than 90 . According to Fig. 6 in [27], an 18% error in the assumed center frequency (70 phase difference between I and Q samples) led to a loss of only 1 dB in contrast performance. Additionally, while narrowband signals are the ideal for DSIQ beamforming, reasonable image
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quality was obtained with fractional bandwidths as high as 55% [27]. Ultimately, the target clinical application for the Sonic Window device dictates the losses in image quality that can be tolerated in exchange for the low cost, compact size, and simple operation that it demands. A consequence of the transmit protection scheme is the lack of focusing on transmit and its associated losses in SNR and penetration. However, as discussed in [35], the Sonic Window is intended for imaging shallow targets, which significantly relaxes the penetration requirement, and all image data is captured in parallel after each transmit event, so a plane wave transmit is appropriate. Furthermore, the protection scheme significantly simplifies the transmitter hardware [27], facilitates a flip-chip interconnect scheme in future versions of the Sonic Window (avoiding SNR losses through cabling) [35], and is critical in enabling the use of a 2-D array, which provides array gain benefits and compensates for slightly worse clutter in azimuth with much better performance in elevation [40]. Similarly, a consequence of DSIQ beamforming is its preference for a smaller signal bandwidth, which corresponds to an increase in C-plane thickness. Our experimental system used a 3.3-MHz center frequency and 30% fractional bandwidth generating a slice thickness of approximately 770 m. However, this is more than adequate to satisfy our initial goal of visualizing a 5-mm-diameter blood vessel in the arm for needle or IV line insertion [40] and slice thickness will decrease as the system center frequency increases in future prototypes. The experimental system described in this paper served well as confirmation of our general design approach, novel technologies, and performance tradeoffs. Several modifications will be necessary to realize the long-term concept of a clinically viable Sonic Window device. The grating lobe and channel loss issues can be addressed through an improved interconnect approach providing smaller transducer element pitch and increased channel count. By designing receive electronics ICs with a channel pitch that corresponds to the transducer element pitch, a flip-chip interconnect scheme is possible, providing a straightforward connection between transducer elements and receive channels with little or no routing [8], [30], [31], [41], [42]. If DSIQ beamforming is used, digitization and memory can be brought on-chip along with the front-end receive electronics, enabling the simultaneous capture of all the complex data needed to form a C-scan image—a simple depth control would vary the timing of the global S/H clock signals. Our experimental system took as long as two minutes to form one C-scan image, which included time for 512 transmits, memory transfer of the volume data from the acquisition card, and processing on the PC. However, with parallel on-chip data capture, DSIQ beamforming could be performed in real-time by a low cost, commercially available DSP [27], [42]. The significantly low pulse repetition frequency (as low as the frame rate) requirements of such an approach can be exploited by switching off the transmit and receive electronics in between transmit events to save power and prevent overheating, even allowing for multiple transmit events for signal averaging to improve SNR [35] and compounding techniques for speckle reduction [27].
VII. CONCLUSION A high-channel-count experimental ultrasound system was constructed and experimentally shown to successfully form pulse-echo images of a wire target and edge phantom. Images formed using the DSIQ beamforming algorithm compared favorably with that of conventional time-delay beamforming and baseband demodulated I/Q beamforming. Image quality was impacted predominantly by grating lobes caused by suboptimal transducer element spatial sampling and routing errors in the transducer PCB coupled with poor surface mount connector contact. Future efforts will focus on reducing element pitch and utilizing alternative interconnect approaches to provide significant improvements in image quality in the future Sonic Window system. ACKNOWLEDGMENT The authors would like to thank E. Girard for her contribution to the transducer design, E. Brush for his assistance in probing the transducer interconnect, and M. Oberhardt for constructing our phantoms. REFERENCES [1] L. D. Greenbaum, “It is time for the sonoscope,” J. Ultrasound Med., vol. 22, no. 4, pp. 321–322, 2003. [2] R. A. Filly, “Is it time for the sonoscope? If so, then let’s do it right!,” J. Ultrasound Med., vol. 22, no. 4, pp. 323–325, 2003. [3] J. R. T. C. Roelandt, “Ultrasound stethoscopy: A renaissance of the physical examination?,” Heart, vol. 89, pp. 971–974, 2003. [4] S. L. P. Langlois, “Portable ultrasound on deployment,” ADF Health, vol. 4, pp. 77–80, Sep. 2003. [5] D. D. Price, S. R. Wilson, and T. G. Murphy, “Trauma ultrasound feasibility during helicopter transport,” J. Air Med., vol. 19, no. 4, pp. 144–146, Oct.–Dec. 2000. [6] J. Hwang, J. Quistgaard, J. Souquet, and L. A. Crum, “Portable ultrasound device for battlefield trauma,” in Proc. IEEE Ultrason. Symp., 1998, vol. 2, pp. 1663–1667. [7] Y. Saijo, S.-I. Nitta, K. Kobayashi, H. Arai, and Y. Nemoto, “Development of an ultra-portable echo device connected to USB port,” Ultrasonics, vol. 42, pp. 699–703, Apr. 2004. [8] T. White, K. Eriksen, and A. Nicoli, “Three-dimensional ultrasonic imaging with a fully populated 128 128 array,” in Proc. 19th Annu. Int. Conf. IEEE Eng. Med. Biol. Soc., 1997, vol. 2, pp. 744–746. [9] J. J. Kim and T. K. Song, “An ultrasound beamforming method using 1.5-bit ADCs for portable ultrasound scanners,” in Proc. IEEE Utrason. Symp., 23–27 Aug. 2004, vol. 3, pp. 1722–1724. [10] M. Karaman and M. Donnell, “Synthetic aperture imaging for small scale systems,” IEEE Trans. Ultrason., Ferroelectr., Freq. Control., vol. 42, no. 3, pp. 429–442, May 1995. [11] R. Fisher, K. Thomenius, R. Wodnicki, R. Thomas, S. Cogan, C. Hazard, W. Lee, D. Mills, B. Khuri-Yakub, and A. Ergun, “Reconfigurable arrays for portable ultrasound,” in Proc. IEEE Ultrason. Symp., 18–21 Sep. 2005, vol. 1, pp. 495–499. [12] J. M. Rothschild, “Ultrasound guidance of central vein catheterization,” ch. 21, 2001 [Online]. Available: http://www.ahrq.gov/clinic/ptsafety/, Evidence Report/Technology Assessment, No. 43. Ch. 21. Making Healthcare Safer. A Critical Analysis of Patient Safety Practices. Agency for Healthcare Research and Quality Publication, No. 01-E058. [13] Y. Beaulieu and P. E. Marik, “Bedside ultrasonography in the ICU: Part 1,” Chest, vol. 128, pp. 881–895, Aug. 2005. [14] P. D. Levin, O. Sheinin, and Y. Gozal, “Use of ultrasound guidance in the insertion of radial artery catheters,” Critical Care Med., vol. 31, pt. 2, pp. 481–484, Feb. 2003. [15] A. G. Randolph, D. J. Cook, C. A. Gonzalez, and C. G. Pribble, “Ultrasound guidance for placement of central venous catheters: A meta-analysis of the literature,” Crit. Care Med., vol. 24, pp. 2053–2058, 1996. [16] W. R. Fry, G. C. Clagett, and P. T. O’Rourke, “Ultrasound-guided central venous access,” Archives Surgery, vol. 134, no. 7, pp. 738–741, 1999. [17] N. Velez, D. E. Earnest, and E. D. Staren, “Diagnostic and interventional ultrasound for breast disease,” Amer. J. Surgery, vol. 180, no. 4, pp. 284–287, 2000.
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[18] A. Liebeskind, A. G. Sikora, A. Komisar, D. Slavit, and K. Fried, “Rates of malignancy in incidentally discovered thyroid nodules evaluated with sonography and fine-needle aspiration,” J. Ultrasound Med, vol. 24, pp. 629–634, 2005. [19] H. T. Harcke, A. D. Levy, and G. J. Lonergan, “The sonographic appearance and detectability of nonopaque and semiopaque materials of military origin,” Military Med., vol. 167, no. 6, pp. 459–463, Jun. 2002. [20] R. Hill, R. Conron, P. Greissinger, and M. Heller, “Ultrasound for the detection of foreign bodies in human tissue,” Ann. Emergency Med., vol. 29, no. 3, pp. 353–356, Mar. 1997. [21] Y. Beaulieu and P. E. Marik, “Bedside ultrasonography in the ICU: Part 2,” Chest, vol. 128, pp. 1766–1781, 2005. [22] J. Ulrich and C. Voit, “Ultrasound in dermatology. Part II. Ultrasound of regional lymph node basins and subcutaneous tumours,” Eur. J. f Dermatol., vol. 11, no. 1, pp. 73–79, 2001. [23] M. L. Ghirardelli, V. Jemos, and P. G. Gobbi, “Diagnostic approach to lymph node enlargement,” Haematologica, vol. 84, pp. 242–247, 1999. [24] M. T. Reeder, B. H. Dick, J. K. Atkins, A. B. Pribis, and J. M. Martinez, “Stress fractures. Current concepts of diagnosis and treatment,” Sports Med., vol. 22, no. 3, pp. 198–212, Sep. 1996. [25] M. I. Fuller, T. N. Blalock, J. A. Hossack, and W. F. Walker, “A portable, low-cost, highly integrated, 3D medical ultrasound system,” in Proc. IEEE Ultrason. Symp., Oct. 2003, vol. 1, pp. 38–41. [26] M. I. Fuller, K. Ranganathan, S. Zhou, T. N. Blalock, J. A. Hossack, and W. F. Walker, “Portable, low-cost medical ultrasound device prototype,” in Proc. IEEE Ultrason. Symp., 23–27 Aug. 2004, vol. 1, pp. 106–109. [27] K. Ranganathan, M. K. Santy, T. N. Blalock, J. A. Hossack, and W. F. Walker, “Direct sampled I/Q beamforming for compact and very lowcost ultrasound imaging,” IEEE Trans. Ultrason., Ferroelectr., Freq. Control., vol. 51, no. 9, pp. 1082–1094, Sep. 2004. [28] E. Girard, S. Zhou, W. Walker, T. Blalock, and J. Hossack, “High element count two dimensional transducer array,” in Proc. IEEE Ultrason. Symp., 5–8 Oct. 2003, vol. 1, pp. 964–967. [29] J. Morizio, S. Guhados, J. Castellucci, and O. von Ramm, “64-channel ultrasound transducer amplifier,” in Proc. Southwest Symp. Mixed-Signal Design, Feb. 2003, pp. 228–232. [30] I. O. Wygant, X. Zhuang, D. T. Yeh, S. Vaithilingam, A. Nikoozadeh, O. Oralkan, A. S. Ergun, M. Karaman, and B. T. Khuri-Yakub, “An endoscopic imaging system based on a two-dimensional CMUT array: Real-time imaging results,” in Proc. IEEE Ultrason. Symp., Sep. 2005, vol. 2, pp. 792–795. [31] I. Çiçek, A. Bozkurt, and M. Karaman, “Design of a front-end integrated circuit for 3D acoustic imaging using 2-D CMUT arrays,” IEEE Trans. Ultrason., Ferreolectr., Freq. Control, vol. 52, no. 12, pp. 2235–2241, Dec. 2005. [32] L. L. Lay, S. J. Carey, and J. V. Hatfield, “Pre-amplifier arrays for intraoral ultrasound probe receiving electronics,” in Proc. IEEE Ultrason. Symp., Aug. 2004, vol. 3, pp. 1753–1756. [33] J. R. Talman and S. L. Garverick, “Integrated-circuit implementation of a matched-cell dynamic focusing architecture for a 5-channel, 50-MHz planar annular array,” in Proc. IEEE Ultrason. Symp., Oct. 2001, vol. 2, pp. 1109–1112. [34] M. Sawan, R. Chebli, and A. Kassem, “Integrated front-end receiver for a portable ultrasonic system,” Analog Integrated Circuits Signal Process., vol. 36, pp. 57–67, 2003. [35] M. I. Fuller, T. N. Blalock, J. A. Hossack, and W. F. Walker, “Novel transmit protection scheme for ultrasound systems,” IEEE Trans. Ultrason., Ferreolectr., Freq. Control, vol. 54, no. 1, pp. 79–86, Jan. 2007. [36] F. J. Harris, “On the use of windows for harmonic analysis with the discrete Fourier Transform,” Proc. IEEE, vol. 66, no. 1, pp. 51–83, Jan. 1978. [37] F. Viola and W. F. Walker, “A spline-based algorithm for continuous time-delay estimation using sampled data,” IEEE Trans. Ultrason., Ferreolectr., Freq. Control, vol. 52, no. 1, pp. 80–93, Jan. 2005. [38] L. A. Negron, F. Viola, E. P. Black, C. A. Toth, and W. F. Walker, “Development and characterization of a vitreous mimicking material for radiation force imaging,” IEEE Trans. Ultrason., Ferreolectr., Freq. Control, vol. 49, no. 11, pp. 1543–1551, Nov. 2002. [39] J. W. Goodman, Introduction to Fourier Optics, 2nd ed. New York: McGraw-Hill, 1988, pp. 46–50. [40] K. Ranganathan and W. F. Walker, “Cystic resolution: A performance metric for ultrasound imaging systems,” IEEE Trans. Ultrason., Ferroelectr., Freq. Control, vol. 54, no. 4, pp. 782–792, Apr. 2007. [41] M. I. Fuller, E. V. Brush, M. D. C. Eames, T. N. Blalock, J. A. Hossack, and W. F. Walker, “The sonic window: Second generation prototype of low-cost, fully integrated, pocket-sized medical ultrasound device,” in Proc. 2005 IEEE Ultrason. Symp., 18–21 Sep. 2005, vol. 1, pp. 273–276.
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Michael I. Fuller (S’01–M’06) received the B.S. and M.S. degrees in electrical engineering from the University of Virginia, Charlottesville, in 2001 and the Ph.D. degree in biomedical engineering, University of Virginia in 2007. After completing his doctoral work, he joined PocketSonics, Inc.—a startup devoted to low-cost and handheld ultrasound systems—where he is currently a Research Scientist and involved in system and integrated circuit development. He is also a Visiting Research Scientist at the University of Virginia. His research interests include analog and mixed-signal integrated circuit design, ultrasound imaging, and medical devices.
Karthik Ranganathan (S’00–M’05) received the B.E. degree in biomedical engineering from the University of Mumbai, Mumbai, India, in 1999 and the Ph.D. degree in biomedical engineering from the University of Virginia, Charlottesville, in 2005. His research interests during his doctoral work included ultrasound beamforming, signal processing, and angular scatter measurement techniques, and his dissertation explored the two extremes—optimal beamforming for the best attainable image quality in state of the art systems and beamforming for very low-cost systems. After his doctoral work, he joined PocketSonics, Inc.—a startup devoted to low-cost and handheld ultrasound systems—where he is currently a Research Scientist and involved in system and beamformer development. He is also a Visiting Research Scientist at the University of Virginia.
Shiwei Zhou (S’03–M’06) was born in Beijing, China, in 1974. He received the B.S. and M.S. degrees in optical-electrical engineering from the Beijing Institute of Technology, Beijing, in 1996 and 1999, respectively, and the Ph.D. degree in biomedical engineering from the University of Virginia, Charlottesville, in 2005. He joined the Biomedical Engineering Department, University of Virginia, in 2000. He is currently a Research Scientist at Philips Research North America, Briarcliff Manor, NY. His research interests are finite-element analysis (FEA) modeling for various ultrasound transducers and applications, new transducer techniques and optimization, high-intensity focused ultrasound, and digital signal processing techniques in ultrasound.
Travis N. Blalock (M’91) received the B.S. and M.S. degrees from the University of Tennessee, Knoxville, in 1985 and 1988, respectively, and the Ph.D. degree from Auburn University, Auburn, AL, in 1991 under the direction of R. Jaeger. In 1991, he joined Agilent Laboratories, Palo Alto, CA, (formerly Hewlett-Packard Labs) where he was involved in several mixed-signal CMOS design efforts, including disk read channel filters, a massively parallel analog cross-correlation processor with integrated image capture, and high-resolution color liquid crystal on silicon microdisplays. He joined the faculty of the Department of Electrical Engineering, University of Virginia, Charlottesville, in 1998. He is currently leading a mixed-signal CMOS research effort at the University of Virginia focusing on integrated signal processing and imaging in medical ultrasound and biotelemetry. Dr. Blalock is a member of Tau Beta Pi and Eta Kappa Nu.
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John A. Hossack (S’90–M’92–SM’02) was born in Glasgow, U.K., in 1964. Hereceived the B.Eng. Hons(I) degree in electrical and electronic engineering from Strathclyde University, Glasgow, in 1986 and the Ph.D. degree in electrical and electronic engineering in 1990. From 1990 to 1992, he was a Postdoctoral Researcher in the E. L. Ginzton Laboratory, Stanford University, Stanford, CA, working under B. A. Auld’s guidance. His research was on modeling of 0:3 and 1:3 piezoelectric composite transducers. In 1992, he joined Acuson, Mountain View, CA, initially working on transducer design. During his time at Acuson, his interests diversified into beamforming and 3-D imaging. Dr. Hossack was made a Fellow of Acuson for “excellence in technical contribution” in 1999. In 2000, he joined the Biomedical Engineering Department, University of Virginia, Charlottesville. His current research interests relate to transducer design, 3-D ultrasound imaging, and contrast microbubble usage in ultrasound for imaging and therapy.
William F. Walker (S’95–M’96) received the B.S.E. and Ph.D. degrees from Duke University, Durham, NC, in 1990 and 1995, respectively. His dissertation explored fundamental limits on the accuracy of adaptive ultrasound imaging. After completing his doctoral work, he stayed on at Duke as an Assistant Research Professor in the Department of Biomedical Engineering. At the same time, he served as a Senior Scientist and President of NovaSon Corporation, Durham. In 1997 he joined the faculty of the Department of Biomedical Engineering, University of Virginia, Charlottesville, being promoted to Associate Professor in 2003. He is an active founder in two ultrasound-based startup companies in Charlottesville: PocketSonics, Inc. and HemoSonics, LLC. His research interests include aperture domain processing, beamforming, angular scatter imaging, tissue elasticity imaging, low-cost system architectures, and time delay and motion estimation.
